Expandable medical device with ductile hinges

ABSTRACT

An expandable tissue supporting device employs ductile hinges at selected points. When expansion forces are applied to the device as a whole, the ductile hinges concentrate expansion stresses and strains in small well defined areas. The expandable medical device including ductile hinges provides the advantages of low expansion force requirements, relatively thick walls which are radio-opaque, improved crimping properties, high crush strength, reduced elastic recoil after implantation, and control of strain to a desired level. The expandable tissue supporting device includes a plurality of elongated beams arranged in a cylindrical device and connected together by a plurality of ductile hinges. The ductile hinges can have a substantially constant hinge cross sectional area which is smaller than a beam cross sectional area such that as the device is expanded from a first diameter to a second diameter, the ductile hinges experience plastic deformation while the beams are not plastically deformed.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a continuation of prior application Ser. No.12/139,737, filed Jun. 16, 2008; which is a continuation of priorapplication Ser. No. 11/837,416, filed Aug. 10, 2007, now abandoned;which is a continuation of application Ser. No. 10/726,605, filed Dec.4, 2003, now U.S. Pat. No. 7,279,004; which is a continuation ofapplication Ser. No. 10/231,007 filed Aug. 30, 2002, now abandoned;which is a continuation of application Ser. No. 09/649,217 filed Aug.28, 2000, now U.S. Pat. No. 6,562,065; which is a continuation ofapplication Ser. No. 09/183,555 filed Oct. 29, 1998, now U.S. Pat. No.6,241,762; which claims priority to Provisional Application No.60/079,881, filed Mar. 30, 1998, now expired, the entire contents ofeach such application incorporated herein by reference.

BACKGROUND OF THE INVENTION

1. Field of the Invention

The present invention relates to tissue-supporting medical devices, andmore particularly to expandable, non-removable devices that areimplanted within a bodily lumen of a living animal or human to supportthe organ and maintain patency.

2. Summary of the Related Art

In the past, permanent or biodegradable devices have been developed forimplantation within a body passageway to maintain patency of thepassageway. These devices are typically introduced percutaneously, andtransported transluminally until positioned at a desired location. Thesedevices are then expanded either mechanically, such as by the expansionof a mandrel or balloon positioned inside the device, or expandthemselves by releasing stored energy upon actuation within the body.Once expanded within the lumen, these devices, called stents, becomeencapsulated within the body tissue and remain a permanent implant.

Known stent designs include monofilament wire coil stents (U.S. Pat. No.4,969,458); welded metal cages (U.S. Pat. Nos. 4,733,665 and 4,776,337);and, most prominently, thin-walled metal cylinders with axial slotsformed around the circumference (U.S. Pat. Nos. 4,733,665, 4,739,762,and 4,776,337). Known construction materials for use in stents includepolymers, organic fabrics and biocompatible metals, such as, stainlesssteel, gold, silver, tantalum, titanium, and shape memory alloys such asNitinol.

U.S. Pat. Nos. 4,733,665, 4,739,762, and 4,776,337 disclose expandableand deformable interluminal vascular grafts in the form of thin-walledtubular members with axial slots allowing the members to be expandedradially outwardly into contact with a body passageway. After insertion,the tubular members are mechanically expanded beyond their elastic limitand thus permanently fixed within the body. The force required to expandthese tubular stents is proportional to the thickness of the wallmaterial in a radial direction. To keep expansion forces withinacceptable levels for use within the body (e.g., 5-10 atm), thesedesigns must use very thin-walled materials (e.g., stainless steeltubing with 0.0025 inch thick walls). However, materials this thin arenot visible on conventional fluoroscopic and x-ray equipment and it istherefore difficult to place the stents accurately or to find andretrieve stents that subsequently become dislodged and lost in thecirculatory system.

Further, many of these thin-walled tubular stent designs employ networksof long, slender struts whose width in a circumferential direction istwo or more times greater than their thickness in a radial direction.When expanded, these struts are frequently unstable, that is, theydisplay a tendency to buckle, with individual struts twisting out ofplane. Excessive protrusion of these twisted struts into the bloodstreamhas been observed to increase turbulence, and thus encourage thrombosis.Additional procedures have often been required to attempt to correctthis problem of buckled struts. For example, after initial stentimplantation is determined to have caused buckling of struts, a second,high-pressure balloon (e.g., 12 to 18 atm) would be used to attempt todrive the twisted struts further into the lumen wall. These secondaryprocedures can be dangerous to the patient due to the risk of collateraldamage to the lumen wall.

Many of the known stents display a large elastic recovery, known in thefield as “recoil,” after expansion inside a lumen. Large recoilnecessitates over-expansion of the stent during implantation to achievethe desired final diameter. Over-expansion is potentially destructive tothe lumen tissue. Known stents of the type described above experiencerecoil of up to about 6 to 12% from maximum expansion.

Large recoil also makes it very difficult to securely crimp most knownstents onto delivery catheter balloons. As a result, slippage of stentson balloons during interlumenal transportation, final positioning, andimplantation has been an ongoing problem. Many ancillary stent securingdevices and techniques have been advanced to attempt to compensate forthis basic design problem. Some of the stent securing devices includecollars and sleeves used to secure the stent onto the balloon.

Another problem with known stent designs is non-uniformity in thegeometry of the expanded stent. Non-uniform expansion can lead tonon-uniform coverage of the lumen wall creating gaps in coverage andinadequate lumen support. Further, over expansion in some regions orcells of the stent can lead to excessive material strain and evenfailure of stent features. This problem is potentially worse in lowexpansion force stents having smaller feature widths and thicknesses inwhich manufacturing variations become proportionately more significant.In addition, a typical delivery catheter for use in expanding a stentincludes a balloon folded into a compact shape for catheter insertion.The balloon is expanded by fluid pressure to unfold the balloon anddeploy the stent. This process of unfolding the balloon causes unevenstresses to be applied to the stent during expansion of the balloon dueto the folds causing the problem non-uniform stent expansion.

U.S. Pat. No. 5,545,210 discloses a thin-walled tubular stentgeometrically similar to those discussed above, but constructed of anickel-titanium shape memory alloy (“Nitinol”). This design permits theuse of cylinders with thicker walls by making use of the lower yieldstress and lower elastic modulus of martensitic phase Nitinol alloys.The expansion force required to expand a Nitinol stent is less than thatof comparable thickness stainless steel stents of a conventional design.However, the “recoil” problem after expansion is significantly greaterwith Nitinol than with other materials. For example, recoil of a typicaldesign Nitinol stent is about 9%. Nitinol is also more expensive, andmore difficult to fabricate and machine than other stent materials, suchas stainless steel.

All of the above stents share a critical design property: in eachdesign, the features that undergo permanent deformation during stentexpansion are prismatic, i.e., the cross sections of these featuresremain constant or change very gradually along their entire activelength. To a first approximation, such features deform under transversestress as simple beams with fixed or guided ends: essentially, thefeatures act as a leaf springs. These leaf spring like structures areideally suited to providing large amounts of elastic deformation beforepermanent deformation commences. This is exactly the opposite of idealstent behavior. Further, the force required to deflect prismatic stentstruts in the circumferential direction during stent expansion isproportional to the square of the width of the strut in thecircumferential direction. Expansion forces thus increase rapidly withstrut width in the above stent designs. Typical expansion pressuresrequired to expand known stents are between about 5 and 10 atmospheres.These forces can cause substantial damage to tissue if misapplied.

FIG. 1 shows a typical prior art “expanding cage” stent design. Thestent 10 includes a series of axial slots 12 formed in a cylindricaltube 14. Each axial row of slots 12 is displaced axially from theadjacent row by approximately half the slot length providing a staggeredslot arrangement. The material between the slots 12 forms a network ofaxial struts 16 joined by short circumferential links 18. The crosssection of each strut 16 remains constant or varies gradually along theentire length of the strut and thus the rectangular moment of inertiaand the elastic and plastic section moduli of the cross section alsoremain constant or vary gradually along the length of the strut. Such astrut 16 is commonly referred to as a prismatic beam. Struts 16 in thistype of design are typically 0.005 to 0.006 inches (0.127-0.1524 mm)wide in the circumferential direction. Strut thicknesses in the radialdirection are typically about 0.0025 inches (0.0635 mm) or less to keepexpansion forces within acceptable levels. However, most stent materialsmust be approximately 0.005 inches (0.127 mm) thick for good visibilityon conventional fluoroscopic equipment. This high ratio of strut widthto thickness, combined with the relatively high strut length and theinitial curvature of the stent tubing combine to cause the instabilityand bucking often seen in this type of stent design. When expanded, thestent structure of FIG. 1 assumes the roughly diamond pattern commonlyseen in expanded sheet metal.

Another stent described in PCT publication number WO 96/29028 usesstruts with relatively weak portions of locally-reduced cross sectionswhich on expansion of the stent act to concentrate deformation at theseareas. However, as discussed above non-uniform expansion is even more ofa problem when smaller feature widths and thicknesses are involvedbecause manufacturing variations become proportionately moresignificant. The locally-reduced cross section portions described inthis document are formed by pairs of circular holes. The shape of thelocally-reduced cross section portions undesirably concentrates theplastic strain at the narrowest portion. This concentration of plasticstrain without any provision for controlling the level of plastic strainmakes the stent highly vulnerable to failure.

In view of the drawbacks of the prior art stents, it would beadvantageous to be able to expand a stent with an expansion force at alow level independent of choice of stent materials, material thickness,or strut dimensions.

It would further be advantageous to have a tissue-supporting device thatpermits a choice of material thickness that could be viewed easily onconventional fluoroscopic equipment for any material.

It would also be advantageous to have a tissue-supporting device that isinherently stable during expansion, thus eliminating buckling andtwisting of structural features during stent deployment.

It would also be desirable to control strain to a desired level whichtakes advantage of work hardening without approaching a level of plasticstrain at which failure may occur.

In addition, it would be advantageous to have a tissue-supporting devicewith minimal elastic recovery, or “recoil” of the device afterexpansion.

It would be advantageous to have a tissue supporting device that can besecurely crimped to the delivery catheter without requiring specialtools, techniques, or ancillary clamping features.

It would further be advantageous to have a tissue-supporting device thathas improved resistance to compressive forces (improved crush strength)after expansion.

It would also be advantageous to have a tissue-supporting device thatachieves all the above improvements with minimal foreshortening of theoverall stent length during expansion.

SUMMARY OF THE INVENTION

The present invention addresses several important problems in expandablemedical device design including: high expansion force requirements; lackof radio-opacity in thin-walled stents; buckling and twisting of stentfeatures during expansion; poor crimping properties; and excessiveelastic recovery (“recoil”) after implantation. The invention alsoprovides benefits of improved resistance to compressive forces afterexpansion, control of the level of plastic strain, and low axialshortening during expansion. Some embodiments of the invention alsoprovide improved uniformity of expansion by limiting a maximum geometricdeflection between struts. The invention may also incorporate sites forthe inclusion of beneficial agent delivery.

The invention involves the incorporation of stress/strain concentrationfeatures or “ductile hinges” at selected points in the body of anexpandable cylindrical medical device. When expansion forces are appliedto the device as a whole, these ductile hinges concentrate expansionstresses and strains in small, well-defined areas while limiting strutdeflection and plastic strain to specified levels.

In accordance with one aspect of the present invention, an expandablemedical device includes a plurality of elongated beams having asubstantially constant beam cross sectional area along a beam length.The plurality of elongated beams are joined together to form asubstantially cylindrical device which is expandable from a cylinderhaving a first diameter to a cylinder having a second diameter. Aplurality of ductile hinges connect the plurality of beams together inthe substantially cylindrical device. The ductile hinges have asubstantially constant hinge cross sectional area along a substantialportion of a hinge length. The hinge cross sectional area is smallerthan the beam cross sectional area such that as the device is expandedfrom the first diameter to the second diameter the ductile hingesexperience plastic deformation while the beams are not plasticallydeformed.

In accordance with a further aspect of the invention, an expandablemedical device includes a cylindrical tube, and a plurality of axialslots formed in the cylindrical tube in a staggered arrangement todefine a network of elongated struts, wherein each of the elongatedstruts are axially displaced from adjacent struts. A plurality ofductile hinges are formed between the elongated struts. The ductilehinges allow the cylindrical tube to be expanded or compressed from afirst diameter to a second diameter by deformation of the ductilehinges. The ductile hinges are asymmetrically configured to reach apredetermined strain level upon a first percentage expansion and toreach the predetermined strain level upon a second percentage ofcompression, wherein the first percentage is larger than the secondpercentage.

In accordance with another aspect of the present invention, anexpandable medical device includes a plurality of elongated beams havinga substantially constant beam cross sectional area along a beam length.A plurality of ductile hinges connect the plurality of beams together ina substantially cylindrical device which is expandable or compressiblefrom a first diameter to a second diameter by plastic deformation of theductile hinges. A plurality of deflection limiting members arepositioned at a plurality of the ductile hinges which limit thedeflection at the ductile hinges.

BRIEF DESCRIPTION OF THE DRAWINGS

The invention will now be described in greater detail with reference tothe preferred embodiments illustrated in the accompanying drawings, inwhich like elements bear like reference numerals, and wherein:

FIG. 1 is an isometric view of a prior art tissue-supporting device;

FIG. 2 is an isometric view of a tissue-supporting device in accordancewith one embodiment of the invention;

FIGS. 3 a-d are perspective views of ductile hinges according to severalvariations of the invention;

FIG. 3 e is a side view of another embodiment of a ductile hinge;

FIGS. 4 a and 4 b are an isometric view and an enlarged side view of atissue-supporting device in accordance with an alternative embodiment ofthe invention;

FIGS. 5 a-c are perspective, side, and cross-sectional views of anidealized ductile hinge for purposes of analysis;

FIG. 5 d is a stress/strain curve for the idealized ductile hinge;

FIG. 6 is a perspective view of a simple beam for purposes ofcalculation;

FIG. 7 is a moment verses curvature graph for a rectangular beam;

FIG. 8 is an enlarged side view of a bent ductile hinge;

FIGS. 9 a and 9 b are enlarged side views of ductile hinges in initialand expanded positions with shortened struts to illustrate axialcontraction relationships;

FIG. 10 is a side view of a portion of an alternative embodiment of atissue supporting device having a high-crush-strength and low-recoil;and

FIG. 11 is an enlarged side view of a tissue-supporting device inaccordance with an alternative embodiment of the invention.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

FIG. 2 shows one embodiment of an expandable tissue supporting device 20in accordance with the present invention. The tissue supporting device20 includes a series of axial slots 22 formed in a cylindrical tube 24.Each axial slot 22 is displaced axially from the slots in adjacent rowsof slots by approximately half the slot length resulting in a staggeredslot arrangement. The offset between adjacent rows of slots results inalternate rows of slots which extend to the ends of the cylindrical tube24. At each interior end of each of the axial slots 22 a circumferentialslot 26 is formed. The material between the slots 22 forms a network ofaxial struts 28 extending substantially parallel to an axis of the tube24. The axial struts 28 are joined by short circumferential links 30.The circumferential links 30 are positioned at both the interior of thecylindrical tube and at the ends of the cylindrical tube. The crosssection (and rectangular moment of inertia) of each of the struts 28 isnot constant along the length of the strut. Rather, the strut crosssection changes abruptly at both ends of each strut 28 at the locationof the circumferential slots 26. The struts 28 are thus not prismatic.Each individual strut 28 is linked to the rest of the structure througha pair of reduced sections 32, one at each end, which act asstress/strain concentration features. The reduced sections 32 of thestruts function as hinges in the cylindrical structure. Since thestress/strain concentration features 32 are designed to operate into theplastic deformation range of generally ductile materials, they arereferred to as ductile hinges. Such features are also commonly referredto as “Notch Hinges” or “Notch Springs” in ultra-precision mechanismdesign, where they are used exclusively in the elastic range.

With reference to the drawings and the discussion, the width of anyfeature is defined as its dimension in the circumferential direction ofthe cylinder. The length of any feature is defined as its dimension inaxial direction of the cylinder. The thickness of any feature is definedas the wall thickness of the cylinder.

The presence of the ductile hinges 32 allows all of the remainingfeatures in the tissue supporting device to be increased in width or thecircumferentially oriented component of their respective rectangularmoments of inertia—thus greatly increasing the strength and rigidity ofthese features. The net result is that elastic, and then plasticdeformation commence and propagate in the ductile hinges 32 before otherstructural elements of the device undergo any significant elasticdeformation. The force required to expand the tissue supporting device20 becomes a function of the geometry of the ductile hinges 32, ratherthan the device structure as a whole, and arbitrarily small expansionforces can be specified by changing hinge geometry for virtually anymaterial wall thickness. In particular, wall thicknesses great enough tobe visible on a fluoroscope can be chosen for any material of interest.

In order to get minimum recoil, the ductile hinges 32 should be designedto operate well into the plastic range of the material, and relativelyhigh local strain-curvatures are developed. When these conditions apply,elastic curvature is a very small fraction of plastic or totalcurvature, and thus when expansion forces are relaxed the percent changein hinge curvature is very small. When incorporated into a strut networkdesigned to take maximum advantage of this effect, the elasticspringback, or “recoil,” of the overall stent structure is minimized.

In the embodiment of FIG. 2, it is desirable to increase the width ofthe individual struts 28 between the ductile hinges 32 to the maximumwidth that is geometrically possible for a given diameter and a givennumber of struts arrayed around that diameter. The only geometriclimitation on strut width is the minimum practical width of the slots 22which is about 0.002 inches (0.0508 mm) for laser machining. Lateralstiffness of the struts 28 increases as the cube of strut width, so thatrelatively small increases in strut width significantly increase strutstiffness. The net result of inserting ductile hinges 32 and increasingstrut width is that the struts 28 no longer act as flexible leafsprings, but act as essentially rigid beams between the ductile hinges.All radial expansion or compression of the cylindrical tissue supportingdevice 20 is accommodated by mechanical strain in the hinge features 32,and yield in the hinge commences at very small overall radial expansionor compression.

Yield in ductile hinges at very low gross radial deflections alsoprovides the superior crimping properties displayed by the ductilehinge-based designs. When a tissue supporting device is crimped onto afolded catheter balloon, very little radial compression of the device ispossible since the initial fit between balloon and device is alreadysnug. Most stents simply rebound elastically after such compression,resulting in very low clamping forces and the attendant tendency for thestent to slip on the balloon. Ductile hinges, however, sustainsignificant plastic deformation even at the low deflections occurringduring crimping onto the balloon, and therefore a device employingductile hinges displays much higher clamping forces. The ductile hingedesigns according to the present invention may be securely crimped ontoa balloon of a delivery catheter by hand or by machine without the needfor auxiliary retaining devices commonly used to hold known stents inplace.

The geometric details of the stress/strain concentration features orductile hinges 32 can be varied greatly to tailor the exact mechanicalexpansion properties to those required in a specific application. Themost obvious and straightforward ductile hinges are formed by slots ornotches with rounded roots, as in FIGS. 3 a and 3 c. Since the laserbeams often used to fabricate these features are themselves round, slotsor notches with circular roots are also among the easiest to fabricate.

FIG. 3 a shows a ductile hinge 36 formed by a pair of opposed circulargrooves 38, 40. According to this embodiment the circumferential slot 26has semicircular ends 38 having a radius of curvature r. Outersemicircular grooves 40 opposed the semicircular ends 38 and also have aradius of curvature r. FIG. 3 c shows another ductile hinge 54 formed bya parabolic groove 56.

Generally, the ductile hinges 36 of the embodiment of FIG. 3 a formedbetween pairs of concave curves 38, 40 have a minimum width along a lineconnecting their respective centers of curvature. When the strutsconnected by the ductile hinge are moved apart or together, plasticdeformation is highly concentrated in a region immediately adjacent tothe plane that bisects the hinge at this narrow point.

For smaller deflection, this very high strain concentration at thebisecting plane is acceptable, and in some cases, useful. For stentcrimping purposed, for example, it is desirable to generate relativelylarge plastic deformations at very small deflection angles.

As a practical matter, however, strut deflection angles for deviceexpansion are often in the 25° to 45° range. At these angles, strain atthe root or bisecting plane of concave ductile hinge features can easilyexceed the 50 to 60% elongation-to-failure of 316L stainless steel, oneof the most ductile stent materials. Deflection limiting features whichwill be described further below limit the geometric deflection ofstruts, but these features do not in themselves affect the propagationpattern of plastic deformation in a given ductile hinge design. Forconcave ductile binges at large bend angles, very high strainconcentrations remain. Scanning electron micrographs have confirmed thisanalysis.

In many engineering applications, it is desirable to limit the amount ofstrain, or “cold-work,” in a material to a specified level in order tooptimize material properties and to assure safe operation. For example,in medical applications it is desirable to limit the amount of cold-workin 316L stainless steel to about 30%. At this level, the strength of thematerial is increased, while the material strain is still well below thefailure range. Ideally, therefore, a safe and effective ductile hingeshould not simply limit gross deflection, but reliably limit materialstrain to a specified level.

FIG. 3 b shows a simple ductile hinge design that allows material strainto be limited to some specified level. The ductile hinge of FIG. 3 b isformed by a rectangular circumferential groove 46 with filleted corners48 on one side of a strut, the opposite side 50 of the strut remainingstraight. The ductile hinges 44 are substantially rectangular sectionsbetween the ends of the groove 46 and the side walls 50.

One of the key concepts in FIG. 3 b is that the ductile hinge 44 has aconstant or substantially constant width along at least a portion of itstotal length. In this configuration, there is no local minimum widthalong the ductile hinge axis, as there is with pairs of concave roots.There is therefore no point concentration of stresses and strains alongthe length of the ductile hinge beam during stent expansion. Inparticular, maximum tensile and compressive strains will be distributedevenly along the upper and lower surfaces of the hinge 44 during stentexpansion. With the gross bend angle limited by mechanical stops, whichare described below in detail, the maximum material strain (at the hingesurfaces) can therefore be reliably limited by adjusting the initiallength of the ductile hinge over which the total elongation isdistributed.

FIG. 3 d shows a ductile hinge 60 in a cylindrical wire 62 forincorporating into a wire-form tissue-supporting device. The ductilehinge 60 is formed by a reduced diameter portion of the wire 62. Again,it is important that the ductile hinge have a substantially constantwidth over a portion of its length in order to provide strain control.Preferably, the ductile hinge is prismatic over a portion of its length.Maximum material strain can be varied by adjusting the hinge length. Theductile hinges of the present invention have a constant or substantiallyconstant width over at least ⅓ of the ductile hinge length, andpreferably over at least ½ of the ductile hinge length.

FIG. 3 e shows an asymmetric ductile hinge 63 that produces differentstrain versus deflection-angle functions in expansion and compression.Each of the ductile hinges 64 is formed between a convex surface 68 anda concave surface 69. The ductile hinge 64 according to a preferredembodiment essentially takes the form of a small, prismatic curved beamhaving a substantially constant cross section. However, a thickness ofthe curved ductile hinge 64 may vary somewhat as long as the ductilehinge width remains constant along a portion of the hinge length. Thewidth of the curved beam is measure along the radius of curvature of thebeam. This small curved beam is oriented such that the smaller concavesurface 69 is placed in tension in the device crimping direction, whilethe larger convex surface 68 of the ductile hinges is placed in tensionin the device expansion direction. Again, there is no local minimumwidth of the ductile hinge 64 along the (curved) ductile hinge axis, andno concentration of material strain. During device expansion tensilestrain will be distributed along the convex surface 68 of the hinge 64and maximum expansion will be limited by the angle of the walls of theconcave notch 69 which provide a geometric deflection limiting feature.Maximum tensile strain can therefore be reliably limited by adjustingthe initial length of the convex arc shaped ductile hinge 64 over whichthe total elongation is distributed.

The ductile hinges illustrated in FIGS. 3 a-e are examples of differentstructures that will function as a stress/strain concentrator. Manyother stress/strain concentrator configurations may also be used as theductile hinges in the present invention. The ductile hinges according tothe present invention generally include an abrupt change in width of astrut that functions to concentrate stresses and strains in the narrowersection of the strut. These ductile hinges also generally includefeatures to limit mechanical deflection of attached struts and featuresto control material strain during large strut deflections. Although theductile hinges have been illustrated in FIG. 2 as positioned at the endsof each of the axial slots 22, they may also be positioned at otherlocations in other designs without departing from the present invention.

An alternative embodiment of a tissue supporting device 80 isillustrated in FIG. 4 a and in the enlarged side view of FIG. 4 b. Thetissue supporting device 80 includes a plurality of cylindrical tubes 82connected by S-shaped bridging elements 84. The bridging elements 84allow the tissue supporting device to bend axially when passing throughthe tortuous path of the vasculature to the deployment site and allowthe device to bend when necessary to match the curvature of a lumen tobe supported. The S-shaped bridging elements 84 provide improved axialflexibility over prior art devices due to the thickness of the elementsin the radial direction which allows the width of the elements to berelatively small without sacrificing radial strength. For example, thewidth of the bridging elements 84 may be about 0.0012-0.0013 inches(0.0305-0.0330 mm). Each of the cylindrical tubes 82 has a plurality ofaxial slots 86 extending from an end surface of the cylindrical tubetoward an opposite end surface. A plurality of axial struts 88 havingductile hinges 90 are formed between the axial slots 86. The ductilehinges 90 are formed by circumferential slots 92 formed at the interiorends of the axial slots 86 and opposed notches 94.

The notches 94 each have two opposed angled walls 96 which function as astop to limit geometric deflection of the ductile hinge, and thus limitmaximum device expansion. As the cylindrical tubes 82 are expanded andbending occurs at the ductile hinges 90, the angled side walls 96 of thenotches 94 move toward each other. Once the opposite side walls 96 of anotch come into contact with each other, they resist further expansionof the particular ductile hinge causing further expansion to occur atother sections of the tissue supporting device. This geometricdeflection limiting feature is particularly useful where unevenexpansion is caused by either variations in the tissue supporting device80 due to manufacturing tolerances or uneven balloon expansion.

The tissue supporting device 20, 80 according to the present inventionmay be formed of any ductile material, such as steel, gold, silver,tantalum, titanium, Nitinol, other shape memory alloys, other metals, oreven some plastics. One preferred method for making the tissuesupporting device 20, 80 involves forming a cylindrical tube and thenlaser cutting the slots 22, 26, 86, 92 and notches 94 into the tube.Alternatively, the tissue supporting device may be formed byelectromachining, chemical etching followed by rolling and welding, orany other known method.

The design and analysis of stress/strain concentration for ductilehinges, and stress/strain concentration features in general, is complex.For example, the stress concentration factor for the simplified ductilehinge geometry of FIG. 3 a can be calculated and is given by thefollowing expression where D is the width of the struts 28, h is theheight of the circular grooves 38, 40, and r is the radius of curvatureof the grooves. For purposes of this example the ratio of h/r is takento be 4. However, other ratios of h/r can also be implementedsuccessfully.

$K = {4.935 - {7.586\left( \frac{2h}{D} \right)} + {0.515\left( \frac{2h}{D} \right)^{2}} + {0.432\left( \frac{2h}{D} \right)^{3}}}$

The stress concentration factors are generally useful only in the linearelastic range. Stress concentration patterns for a number of othergeometries can be determined through photoelastic measurements and otherexperimental methods. Stent designs based on the use of stress/strainconcentration features, or ductile hinges, generally involve morecomplex hinge geometries and operate in the non-linear elastic andplastic deformation regimes.

The general nature of the relationship among applied forces, materialproperties, and ductile hinge geometry can be more easily understoodthrough analysis of an idealized hinge 66 as shown in FIGS. 5 a-5 e. Thehinge 66 is a simple beam of rectangular cross section having a width h,length L and thickness b. The idealized hinge 66 haselastic-ideally-plastic material properties which are characterized bythe ideal stress/strain curve of FIG. 5 d. It can be shown that the“plastic” or “ultimate bending moment” for such a beam is given by theexpression:

${M_{p} \equiv M_{ult}} = {\delta_{yp}\frac{{bh}^{2}}{4}}$

Where b corresponds to the cylindrical tube wall thickness, h is thecircumferential width of the ductile hinge, and δ_(yp) is the yieldstress of the hinge material. Assuming only that expansion pressure isproportional to the plastic moment, it can be seen that the requiredexpansion pressure to expand the tissue supporting device increaseslinearly with wall thickness b and as the square of ductile hinge widthh. It is thus possible to compensate for relatively large changes inwall thickness b with relatively small changes in hinge width h. Whilethe above idealized case is only approximate, empirical measurements ofexpansion forces for different hinge widths in several different ductilehinge geometries have confirmed the general form of this relationship.Accordingly, for different ductile hinge geometries it is possible toincrease the thickness of the tissue supporting device to achieveradiopacity while compensating for the increased thickness with a muchsmaller decrease in hinge width.

Ideally, the stent wall thickness b should be as thin as possible whilestill providing good visibility on a fluoroscope. For most stentmaterials, including stainless steel, this would suggest a thickness ofabout 0.005-0.007 inches (0.127-0.178 mm) or greater. The inclusion ofductile hinges in a stent design can lower expansion forces/pressures tovery low levels for any material thickness of interest. Thus ductilehinges allow the construction of optimal wall thickness tissuesupporting devices at expansion force levels significantly lower thancurrent non-visible designs.

The expansion forces required to expand the tissue supporting device 20according to the present invention from an initial condition illustratedin FIG. 2 to an expanded condition is between 1 and 5 atmospheres,preferably between 2 and 3 atmospheres. The expansion may be performedin a known manner, such as by inflation of a balloon or by a mandrel.The tissue supporting device 20 in the expanded condition has a diameterwhich is preferably up to three times the diameter of the device in theinitial unexpanded condition.

Many tissue supporting devices fashioned from cylindrical tubes comprisenetworks of long, narrow, prismatic beams of essentially rectangularcross section as shown in FIG. 6. These beams which makeup the knowntissue supporting devices may be straight or curved, depending on theparticular design. Known expandable tissue supporting devices have atypical wall thickness b of 0.0025 inches (0.0635 mm), and a typicalstrut width h of 0.005 to 0.006 inches (0.127-0.1524 mm). The ratio ofb:h for most known designs is 1:2 or lower. As b decreases and as thebeam length L increases, the beam is increasingly likely to respond toan applied bending moment M by buckling, and many designs of the priorart have displayed this behavior. This can be seen in the followingexpression for the “critical buckling moment” for the beam of FIG. 6.

$M_{crit} = \frac{\pi \; b^{3}h\sqrt{{EG}\left( {1 - {0.63\mspace{14mu} {b/h}}} \right)}}{6L}$

-   -   Where: E=Modulus of Elasticity        -   G=Shear Modulus

By contrast, in a ductile hinge based design according to the presentinvention, only the hinge itself deforms during expansion. The typicalductile hinge 32 is not a long narrow beam as are the struts in theknown stents. Wall thickness of the present invention may be increasedto 0.005 inches (0.127 mm) or greater, while hinge width is typically0.002-0.003 inches (0.0508-0.0762 mm), preferably 0.0025 inches (0.0635mm) or less. Typical hinge length, at 0.002 to 0.005 inches(0.0508-0.0127 mm), is more than an order of magnitude less than typicalstrut length. Thus, the ratio of b:h in a typical ductile hinge 32 is2:1 or greater. This is an inherently stable ratio, meaning that theplastic moment for such a ductile hinge beam is much lower than thecritical buckling moment M_(crit), and the ductile hinge beam deformsthrough normal strain-curvature. Ductile hinges 32 are thus notvulnerable to buckling when subjected to bending moments duringexpansion of the tissue supporting device 20.

To provide optimal recoil and crush-strength properties, it is desirableto design the ductile hinges so that relatively large strains, and thuslarge curvatures, are imparted to the hinge during expansion of thetissue supporting device. Curvature is defined as the reciprocal of theradius of curvature of the neutral axis of a beam in pure bending. Alarger curvature during expansion results in the elastic curvature ofthe hinge being a small fraction of the total hinge curvature. Thus, thegross elastic recoil of the tissue supporting device is a small fractionof the total change in circumference. It is generally possible to dothis because common stent materials, such as 316L Stainless Steel havevery large elongations-to-failure (i.e., they are very ductile).

It is not practical to derive exact expressions for residual curvaturesfor complex hinge geometries and real materials (i.e., materials withnon-idealized stress/strain curves). The general nature of residualcurvatures and recoil of a ductile hinge may be understood by examiningthe moment-curvature relationship for the elastic-ideally-plasticrectangular hinge 66 shown in FIGS. 5 a-c. It may be shown that therelationship between the applied moment and the resulting beam curvatureis:

$M = {{M_{p}\left\lbrack {1 - {{1/3}\left( \frac{y_{o}}{\left. {h/2} \right)} \right)^{2}}} \right\rbrack} = {{3/2}{M_{yp}\left\lbrack {1 - {{1/3}\left( \frac{\kappa_{yp}}{\kappa} \right)^{2}}} \right\rbrack}}}$

This function is plotted in FIG. 7. It may be seen in this plot that theapplied moment M asymptotically approaches a limiting value M_(p),called the plastic or ultimate moment. Beyond 11/12 M_(p), large plasticdeformations occur with little additional increase in applied moment.When the applied moment is removed, the beam rebounds elastically alonga line such as a-b. Thus, the elastic portion of the total curvatureapproaches a limit of 3/2 the curvature at the yield point. Theserelations may be expressed as follows:

$M_{p} = {{{\frac{3}{2}M_{yp}}->\kappa_{rebound}} = {\frac{3}{2}\kappa_{yp}}}$

Imparting additional curvature in the plastic zone cannot furtherincrease the elastic curvature, but will decrease the ratio of elasticto plastic curvature. Thus, additional curvature or larger expansion ofthe tissue supporting device will reduce the percentage recoil of theoverall stent structure.

As shown in FIG. 8, when a rigid strut 28 is linked to the ductile hinge66 described above, the strut 28 forms an angle θ with the horizontalthat is a function of hinge curvature. A change in hinge curvatureresults in a corresponding change in this angle θ. The angular elasticrebound of the hinge is the change in angle Δθ that results from therebound in elastic curvature described above, and thus angular reboundalso approaches a limiting value as plastic deformation proceeds. Thefollowing expression gives the limiting value of angular elastic reboundfor the idealized hinge of FIG. 8.

$\theta_{rebound} = {3 \in_{yp}\frac{L}{h}}$

Where strain at the yield point is an independent material property(yield stress divided by elastic modulus); L is the length of theductile hinge; and h is the width of the hinge. For non-idealizedductile hinges made of real materials, the constant 3 in the aboveexpression is replaced by a slowly rising function of total strain, butthe effect of geometry would remain the same. Specifically, the elasticrebound angle of a ductile hinge decreases as the hinge width hincreases, and increases as the hinge length L increases. To minimizerecoil, therefore, hinge width h should be increased and length L shouldbe decreased.

Ductile hinge width h will generally be determined by expansion forcecriteria, so it is important to reduce hinge length to a practicalminimum in order to minimize elastic rebound. Empirical data on recoilfor ductile hinges of different lengths show significantly lower recoilfor shorter hinge lengths, in good agreement with the above analysis.

The ductile hinges 32 of the tissue supporting device 20 provide asecond important advantage in minimizing device recoil. The embodimentof FIG. 2 shows a network of struts joined together through ductilehinges to form a cylinder. In this design, the struts 28 are initiallyparallel to an axis of the device. As the device is expanded, curvatureis imparted to the hinges 32, and the struts 28 assume an angle θ withrespect to their original orientation, as shown in FIG. 8. The totalcircumferential expansion of the tissue supporting device structure is afunction of hinge curvature (strut angle) and strut length. Moreover,the incremental contribution to stent expansion (or recoil) for anindividual strut depends on the instantaneous strut angle. Specifically,for an incremental change in strut angle Δθ, the incremental change incircumference ΔC will depend on the strut length R and the cosine of thestrut angle θ.

ΔC=RΔθ cos θ

Since elastic rebound of hinge curvature is nearly constant at any grosscurvature, the net contribution to circumferential recoil ΔC is lower athigher strut angles θ. The final device circumference is usuallyspecified as some fixed value, so decreasing overall strut length canincrease the final strut angle θ. Total stent recoil can thus beminimized with ductile hinges by using shorter struts and higher hingecurvatures when expanded.

Empirical measurements have shown that tissue supporting device designsbased on ductile hinges, such as the embodiment of FIG. 2, displaysuperior resistance to compressive forces once expanded despite theirvery low expansion force. This asymmetry between compressive andexpansion forces may be due to a combination of factors including thegeometry of the ductile hinge, the increased wall thickness, andincreased work hardening due to higher strain levels.

According to one example of the tissue supporting device of theinvention, the device can be expanded by application of an internalpressure of about 2 atmospheres or less, and once expanded to a diameterbetween 2 and 3 times the initial diameter can withstand a compressiveforce of about 16 to 20 gm/mm or greater. Examples of typicalcompression force values for prior art devices are 3.8 to 4.0 gm/mm.

While both recoil and crush strength properties of tissue supportingdevices can be improved by use of ductile hinges with large curvaturesin the expanded configuration, care must be taken not to exceed anacceptable maximum strain level for the material being used. For theductile hinge 44 of FIG. 3 b, for example, it may be shown that themaximum material strain for a given bend angle is given by theexpression:

$ɛ_{\max} = {\frac{h}{L}\frac{\theta}{2}}$

Where ε_(max) is maximum strain, h is ductile hinge width, L is ductilehinge length and θ is bend angle in radians. When strain, hinge widthand bend angle are determined through other criteria, this expressioncan be evaluated to determine the correct ductile hinge length L.

For example, suppose the ductile hinge 44 of FIG. 3 b was to befabricated of 316L stainless steel with a maximum strain of 30%; ductilehinge width h is set at 0.0025 inch (0.0635 mm) by expansion forcecriteria; and the bend angle θ is mechanically limited to 0.5 radians(≅30%) at full stent expansion. Solving the above expression for L givesthe required ductile hinge length of at least about 0.0033 inches(0.0838 mm).

Similar expressions may be developed to determine required lengths formore complicated ductile hinge geometries, such as shown in FIG. 3 e.Typical values for the prismatic portions of these curved ductile hingesrange from about 0.002 to about 0.0035 inches (0.051-0.089 mm) in hingewidth and about 0.002 to about 0.006 inches (0.051-0.152 mm) in hingelength. The tissue supporting device design of FIGS. 4 a and 4 b includea stop which limits the maximum geometric deflection at the ductilehinges by the design of the angled walls 96 of the notches 94.

In many designs of the prior art, circumferential expansion wasaccompanied by a significant contraction of the axial length of thestent which may be up to 15% of the initial device length. Excessiveaxial contraction can cause a number of problems in device deploymentand performance including difficulty in proper placement and tissuedamage. Designs based on ductile hinges 32 can minimize the axialcontraction, or foreshortening, of a tissue supporting device duringexpansion as follows.

FIGS. 9 a and 9 b illustrate an exaggerated ductile hinge 32 andshortened struts 28 in initial and expanded conditions. Each strut 28 isattached to two ductile hinges 32 at opposite ends. Each ductile hinge32 has an instant center of rotation C₁, C₂ that is an effective pivotpoint for the attached strut 28. Initially, during expansion the pivotpoint C₁ is displaced vertically by a distance d until C₁ is positionedeven with C₂ as shown in FIG. 9 b. When the array is expandedvertically, the axial struts 28 move in a circular arc with respect tothe pivot points, as shown in FIG. 9 b. It can be seen that thehorizontal distance e between pivot points C₁ and C₂ actually increasesinitially, reaching a maximum e_(max) when the two points are on thesame horizontal axis as shown in FIG. 9 b. As the vertical expansioncontinues, the device compresses axially back to its original length.Only when vertical expansion of the array continues beyond the pointwhere the horizontal distance e between C₁ an C₂ is the same as theoriginal horizontal distance e does the overall length of the arrayactually begin to contract. For the stent shown in FIG. 2, for example,approximately ⅓ of the total circumferential expansion has beenaccomplished by the time the configuration of FIG. 9 b is reached, andthe stent exhibits very low axial contraction.

This ability to control axial contraction based on hinge and strutdesign provides great design flexibility when using ductile hinges. Forexample, a stent could be designed with zero axial contraction.

An alternative embodiment that illustrates the trade off between crushstrength and axial contraction is shown in FIG. 10. FIG. 10 shows aportion of a tissue supporting device 70 having an array of struts 72and ductile hinges 74 in the unexpanded state. The struts 72 arepositioned initially at an angle θ₁ with respect to a longitudinal axisX of the device. As the device is expanded radially from the unexpandedstate illustrated in FIG. 10, the angle θ₁ increases. In this case thedevice contracts axially from the onset of vertical expansion throughoutthe expansion. Once the device bas been completely expanded the finalangle θ₁ made by the strut 72 with the horizontal will be much greaterthan the angle θ in the device of FIGS. 8 a and 8 b. As shownpreviously, a higher final strut angle θ₁, can significantly increasecrush strength and decrease circumferential recoil of the stentstructure. However, there is a trade off between increased crushstrength and increase in axial contraction.

According to one example of the present invention, the struts 72 arepositioned initially at an angle of about 0° to 45° with respect to alongitudinal axis of the device. As the device is expanded radially fromthe unexpanded state illustrated in FIG. 10 a, the strut angle increasesto about 20° to 80°.

According to one alternative embodiment of the present invention, theexpandable tissue supporting device can also be used as a deliverydevice for certain beneficial agents including drugs, chemotherapy, orother agents. Due to the structure of the tissue supporting deviceincorporating ductile hinges, the widths of the struts can besubstantially larger than the struts of the prior art devices. Thestruts due to their large size can be used for beneficial agent deliveryby providing beneficial agent on the struts or within the struts.Examples of beneficial agent delivery mechanisms include coatings on thestruts, such as polymer coatings containing beneficial agents, laserdrilled holes in the struts containing beneficial agent, and the like.

Referring to FIG. 11, an alternative embodiment of a tissue supportingdevice is shown generally by reference number 180, with like referencenumerals being used to denote like parts to those discussed above withrespect to FIG. 4 b. In addition, device 180 includes laser drilledholes 182 in the elongated beams or struts 88 for containing abeneficial agent 184.

While the invention has been described in detail with reference to thepreferred embodiments thereof, it will be apparent to one skilled in theart that various changes and modifications can be made and equivalentsemployed, without departing from the present invention.

1-26. (canceled)
 27. An expandable medical device comprising: aplurality of elongated beams, the plurality of elongated beams joinedtogether to form a substantially cylindrical device which is expandablefrom a first diameter to a second diameter and; a plurality of ductilehinges connecting the plurality of beams together in the substantiallycylindrical device, wherein the hinge width is smaller than the beamwidth such that as the device is expanded from the first diameter to thesecond diameter the ductile hinges experience plastic deformation whilethe beams are not plastically deformed, each of the ductile hinges beingin the shape of a curved beam having first and second arcuate surfacesfacing the same direction with the second arcuate surface being largerthan the first arcuate surface, the curved beams being positioned suchthat during expansion tensile strain is distributed along the secondarcuate surface of the curved prismatic beam.
 28. The expandable medicaldevice according to claim 1, wherein the hinge width is smaller than thehinge thickness.
 29. The expandable medical device according to claim 1,wherein the hinge width is no greater than 60% of the hinge thickness.30. The expandable medical device according to claim 27, wherein thehinge width is at least 50% smaller than the beam width.
 31. Theexpandable medical device according to claim 27, wherein the device isexpandable by a balloon catheter pressurized by an inflation pressure of1 to 5 atmospheres.
 32. The expandable medical device according to claim27, wherein the hinge width is less than ⅔ the beam width.
 33. Theexpandable medical device according to claim 27, wherein a transitionbetween the cross sectional area of the struts and the cross sectionalarea of the ductile hinges is an abrupt transition which extends lessthan 10 percent of a length of a strut.
 34. The expandable medicaldevice according to claim 27, wherein a ratio of a length of the ductilehinges to a length of the axial struts is 1:6 or less.
 35. Theexpandable medical device according to claim 27, wherein the elongatedstruts include a beneficial agent for delivery to a patient.
 36. Theexpandable medical device according to claim 27, wherein ductile hingesare configured such that during crimping of the device onto a balloon,tensile strain is distributed along the first arcuate surface of thecurved prismatic beam.
 37. An expandable medical device comprising: aplurality of elongated beams each defining a beam width in thecircumferential direction of the cylindrical device, the plurality ofelongated beams joined together to form a substantially cylindricaldevice which is expandable from a first diameter to a second diameter;and a plurality of ductile hinges connecting the plurality of beamstogether in the substantially cylindrical device, each hinge defining ahinge width in the circumferential direction of the cylindrical deviceand having first and second side surfaces, wherein the hinge width issmaller than the beam width such that as the device is expanded from thefirst diameter to the second diameter the ductile hinges experienceplastic deformation while the beams are not plastically deformed, theductile hinges being asymmetrically configured with a first side surfaceplaced in compression during expansion of the device and a second sidesurface placed intension during expansion of the device, wherein thefirst side surface has a length smaller than a length of the second sidesurface; wherein the second side surface is a convex arcuate surface.38. The expandable medical device according to claim 37, wherein thehinge width is at least 50% smaller than the beam width.
 39. Theexpandable medical device according to claim 37, wherein the hinge widthis less than ⅔ the beam width.
 40. The expandable medical deviceaccording to claim 37, wherein a transition between the cross sectionalarea of the struts and the cross sectional area of the ductile hinges isan abrupt transition which extends less than 10 percent of a length of astrut.
 41. The expandable medical device according to claim 37, whereinthe elongated struts include a beneficial agent for delivery to apatient.